Optical Coherence Tomography (OCT) is a depth-resolved high-resolution imaging technique. For a typical OCT system, a low-coherence light is illuminated on the sample, the backscattered light from sample is interfered with a reference light, depth-resolved image can be reconstructed based on the interference fringe signal. So far, OCT has been applied mostly for ophthalmology, cardiology, and gastrointestinal (GI) tract imaging. It was first introduced as a time-domain system, which has a mechanically scanning reference mirror to acquire depth reflectivity information of sample, later spectrometer based spectral-domain OCT was invented, fringe signal is acquired from line-scan camera after grating spreads the spectrum, and more recently, swept-source laser based OCT has been developed, spectrum signal is encoded in time series, which substantially increases the acquisition speed.
The OCT systems and methods provide a high-resolution real-time imaging technique: the axial resolution can be around 10 μm or 1 μm in tissue depending on the bandwidth of light source. By applying a 2-dimensional (2D) galvanometer scanning system or rotation-pullback scheme, 3D image can be constructed. Lateral resolution of the image depends on objective optics. The resolution and depth of focus of a conventional objective obeys Rayleigh criteria, for example, using an objective with 0.02 numerical aperture (NA) imaging at 1.3 μm center wavelength, its lateral resolution is approximately 40 μm, and depth of focus 2.0 mm; increasing objective NA leads to higher resolution, but reduces the depth of focus (depth of focus is proportional to the square of the resolution). Therefore, there is always a trade-off between lateral resolution and depth-of-focus.
For cellular or sub-cellular level imaging, the resolution of imaging system requires to be better than 5 μm, less than 2 μm or even less than 1 μm in tissue. By applying an ultra-broad band light source (300 nm spectrum bandwidth), the axial resolution of OCT system can be less than 2 μm in tissue; combined with a 0.1 NA objective imaging at 800 nm wavelength region, the lateral resolution can achieve 5 μm, but with depth-of-focus only 50 μm. So limitation on depth-of-focus prevents the application of high-resolution OCT when long ranging depth or imaging depth is required such as in-vivo cardiology imaging, which requires ranging depths longer than 500 μm, and preferably longer than 1.0 mm, and in best cases, longer than 1.5 mm. In addition, for many arterial applications, it cannot be assured that the catheter is in the center of the vessel, which places the desired specifications to be at least greater than 2.0 mm and preferably greater than 3.0 mm and in the best case, greater than 5.0 mm. Previously, techniques such as aperture apodization and synthesized aperture have been developed to increase the depth of focus, but application of these techniques for in-vivo imaging still presents difficulty due to limited performance and implementation complexity.
In a depth-resolved imaging system (such as optical coherence tomography (OCT)), trade-off usually exists between lateral resolution and depth of focus (DOF). For instance, the DOF of a Gaussian beam is proportional to the square of the size of the focal spot, therefore using a Gaussian beam to acquire images that can maintain a high lateral resolution over a long axial field of view is challenging. To overcome this obstacle, various techniques have been proposed, and can be divided into four categories: 1. numerical refocusing-a method that digitally compensates the defocus aberration of the beam according to beam diffraction and sample scattering, which presents difficulty in real-time imaging due to computational intensive processing and phase-stable acquisition; 2. multi-beam acquisition scheme-a technique that applies multiple beams to acquire in-focus images from different depth regions of the sample, which substantially increases the complexity of the imaging system; 3. pupil apodization and phase mask-methods that modify the optical system pupil to extend the DOF of a standard beam, which provides limited improvement while the signal loss is critical; 4. application of a Bessel beam-one type of the beam that has depth-invariant Bessel profile on the transverse plane able to provide an extremely long DOF.
In a coherent light based optical system, the most common form of beam is a Gaussian beam as it is the output of a single-mode fiber, and after transmitted by the standard optical components such as spherical lenses, reflective mirrors, and spacers, it remains as a Gaussian beam, so we define this type of optical system as a Gaussian beam optical system. A Bessel beam can be generated in a Gaussian beam optical system by using an axicon lens. If a Gaussian beam is focused by an axicon lens, the energy of the focused field is concentrated at the tip of the axicon lens, and falls off quickly in the axial direction in response to transverse Gaussian intensity distribution of the input beam. Extension of the working distance requires more complicated optical setup such as introducing a hollow beam that has a lower intensity distribution in the center of the beam, which unavoidably leads to increased system complexity and makes the design not suitable for implementation in a compact imaging device such as a catheter or an endoscope, of which working distance is an important parameter that needs to meet specific application specification, such as 300-500 μm for human coronary artery catheter and 6-7 mm for human gastrointestinal (GI) tract capsule. Thus, the application of a Bessel beam for DOF extension in a catheter or endoscope optical system is greatly limited.
Accordingly, there is a need to address and/or overcome at least some of the deficiencies described herein above.